C-arm x-ray devices, i.e. x-ray devices having a C-arm and on which an x-ray source and an x-ray detector are provided opposite one another are already known in the prior art. As a result of at least one degree of freedom of movement of the C-arm, at least one rotation facility, the x-ray detector and the x-ray source can be moved around an object to be imaged, such as a patient, so that images can be recorded from different projection angles, i.e. from different projection geometries.
To avoid influences from scattered rays, these types of x-ray device mostly have what is referred to as an anti scatter grid, which ultimately consists of a grid placed on the x-ray detector. Focused anti scatter grids are used in such cases, which are permanently connected to the x-ray detector and are set to a usual distance between the x-ray source and the x-ray detector, as are needed for explicit, such as high-resolution, two-dimensional diagnostic x-ray images, known as radiography images.
In the interim however it has also been proposed that three-dimensional images be recorded by an x-ray device with a C-arm, by a method similar to that of computed tomography (CT) being used, giving rise to the frequently used name C-arm CT. In such cases two-dimensional projection images are mostly recorded along a recording trajectory, mostly by rotation of the C-arm, from different recording geometries including different projection directions, from which three-dimensional image datasets can then be constructed using usual methods, for example by iterative reconstruction or filtered back projection (FBP). The three-dimensional image dataset in such cases ultimately contains a three-dimensional volume of the attenuation coefficients which describes the target.
The said anti scatter grid not only leads to a reduction in scattered radiation but also in the desired primary radiation since it has a structure, such as a grid structure, also partly blocking the primary radiation. This leads to artifacts in the x-ray projection image, which in their turn lead to a reduced image quality in the reconstructed image dataset. In order to reduce these artifacts it is known that pixel-based gain correction can be undertaken by a correction image, which for example can be obtained by calibration. In such cases not just the effect of the grid structures of the anti scatter grid can be taken into account by such a correction image, but also a problem which frequently arises precisely with C-arm CT, since a greater distance between the x-ray source and the x-ray detector is used than that for which the anti scatter grid is focused, so that in the final analysis the images are recorded with a defocused anti scatter grid, which leads to an unequal intensity distribution, to a reduction in intensity at the edges of the detector, even with homogeneous illumination.
In conjunction with such an intensity reduction, as a result of a defocused anti scatter grid, a further disadvantageous effect is also to be observed in C-arm CT, namely the fact that, because of mechanical inaccuracies of the x-ray device, deviations occur at different recording times, i.e. with different recording geometries, in respect of the position of the focus of the x-ray source relative to the x-ray detector. This also results in displacement of the artifact patterns and also to the said intensity distribution, which can lead to clearly visible edges in the reconstructed volume, although this effect is barely perceptible in the projection images. These types of mechanical inaccuracies of the x-ray device are based on the fact that the C-arm is not rigid and thus for example, in different positions of the C-arm, a different gravity influences the x-ray device in different ways.
The said mechanical inaccuracies ultimately do not allow a global correction, including a single correction image, to be used for all recording geometries, since this in turn can lead to the retention of artifacts or to the occurrence of new artifacts, as described.
To resolve this problem it has been proposed that a separate pixel-based correction image be created and kept for each recording geometry, including each projection image to be recorded. The correction image is determined by a calibration measurement, wherein for example, to arrive at similar intensity values as during the imaging of a patient, a copper plate can be introduced into the radiation path. Since noise is a critical parameter in the reconstruction, since for example the noise is directly correlated to patients with the patient dose, the correction image may only contain a very small noise amount. In the prior art, to obtain good correction images, a plurality of calibration images is recorded and averaged for each recording geometry, which demands a high calibration effort. Removing noise from the images by simple smoothing, for example a Gaussian filter, does not resolve the problem since the pattern of the artifacts on the correction images is also destroyed by this, so that the correction can no longer rectify the destroyed artifact pattern. A further disadvantage of the known method is that a very high amount of memory space is required to store correction images for each imaging geometry, since the significant structures would also be lost within the framework of a compression.